Positron Emission Tomography (PET) has gained significant popularity in nuclear medicine because of the ability to non-invasively study physiological processes within the body. Applications employing the PET technology for its sensitivity and accuracy include those in the fields of oncology, cardiology and neurology.
Using compounds such as .sup.11 C-labeled glucose, .sup.18 F-labeled glucose, .sup.13 N-labeled ammonia and .sup.15 O-labeled water, PET can be used to study such physiological phenomena as blood flow, tissue viability, and in vivo brain neuron activity. Positrons emitted by these neutron deficient compounds interact with free electrons in the body area of interest, resulting in the annihilation of the positron. This annihilation yields the simultaneous emission of a pair of photons (gamma rays) approximately 180.degree. (angular) apart. A compound having the desired physiological effect is administered to the patient, and the radiation resulting from annihilation is detected by a PET tomograph. After acquiring these annihilation "event pairs" for a period of time, the isotope distribution in a cross section of the body can be reconstructed.
PET data acquisition occurs by detection of both photons emitted from the annihilation of the positron in a coincidence scheme. Due to the approximate 180.degree. angle of departure from the annihilation site, the location of the two detectors registering the "event" define a chord passing through the location of the annihilation. By histogramming these lines of response (the chords), a "sinogram" is produced that may be used by a process of back-projection to produce a three dimensional image of the activity. Detection of these lines of activity is performed by a coincidence detection scheme. A valid event line is registered if both photons of an annihilation are detected within a coincidence window of time. Coincidence detection methods ensure (disregarding other second-order effects) that an event line is histogrammed only if both photons originate from the same positron annihilation.
In the traditional (2-D) acquisition of a modern PET tomograph, a collimator (usually tungsten) known as a septa is placed between the object within the field-of-view and the discrete axial rings of detectors. This septa limits the axial angle at which a gamma ray can impinge on a detector, typically limiting the number of axial rings of detectors that a given detector in a specific ring can form a coincidence with to a few rings toward the front of the tomograph from the given detector's ring, the same ring that the detector is within, and a few rings toward the rear of the tomograph from the given detector's ring.
Attenuation was first measured in PET by using a ring of positron emitting isotope surrounding the object to be measured. In this technique, the ratio between a transmission scan and a blank scan form the attenuation. The blank is measured by simply measuring the rate that gamma rays from positrons are detected by the detection system when no attenuating media is present. In the original scanners as described above as having septa, the septa are provided for collimating the gamma rays in an axial direction, but the rings allow for no transaxial collimation. The lack of collimation allow the acceptance of scattered events into the transmission measurement, resulting in an underestimate of the attenuation. To improve the transmission measurement, systems use rotating rod sources. These sources are disposed in parallel fashion to the axis of the scanner and are collimated in the axial direction by the septa. In the transaxial direction, the collimation may be provided electronically since the position of the source is known. However, the activity in the rod must be the same as that activity in the earlier ring source to provide the same count rate. With modern block detectors, the dead-time of the near block limits the activity in the rod.
A more recent advancement in PET acquisition is 3-D, in which the septa are removed, which allows a given detector to be in coincidence with detectors from all other detector rings. With the advent of three-dimensional reconstruction techniques, greater sensitivity to emission counts is possible if the septa are removed. As the septa represent a significant cost, there is also an economic incentive to exclude them from the system. However, with the absence of septa, the problems of both detector dead-time and scatter are magnified.
Since the position of a source with respect to the detector system can be known, there is no need to detect coincidences, thereby allowing the use of a source that emits single gamma rays. Only one detector--the detector on the far side of the system--is needed to make the transmission or blank measurements. Without the counting losses due to the dead-time of the near detector, the activity of the source may be increased resulting in an increase in count-rate and thus a better quality measurement. However, without axial collimation, the scatter included in the transmission scan causes an underestimate of the attenuation measurement. To decrease the possibility of scatter, the gamma rays from the source can be collimated with lead or tungsten to form a beam that illuminates only a narrow plane of detectors. Other gamma rays that would only contribute to background are eliminated. Since the directionality of single gamma rays cannot be determined, only a single point of activity illuminating a detector bank can be used. This requires increased levels of activity to meet the count-rate needed for an adequate quality measurement. Also, the scanning protocol is more efficient if the transmission measurement is performed after the patient has been injected with radioactivity. Even though a different isotope such as .sup.137 Cs which emits gamma rays with an energy of 662 keV can be used for the transmission scan, there is a significant difficulty in distinguishing the transmission events from the emission events.
Another tomographic diagnostic system that is similar to PET is known as single photon emission computed tomography (SPECT). The distinction is that in SPECT, only a single photon from a nuclear decay within the patient is detected. Also, the line of response traveled by the photon is determined exclusively by detector collimation in SPECT, as opposed to the coincident detection of two collinear photons as in PET.
In computed axial tomography (CAT, or now also referred to as CT), an external x-ray source is caused to be passed around a patient. Detectors around the patient then respond to x-ray transmission through the patient to produce an image of an area of study. Unlike PET and SPECT, which are emission tomography techniques because they rely on detecting radiation emitted from the patient, CT is a transmission tomography technique which utilizes only a radiation source external to the patient.
The details of carrying out a PET study are given in numerous publications. Typically, the following references provide a background for PET. These are incorporated herein by reference for any of their teachings.
1. M. E. Phelps, et al.: "Positron Emission Tomography and Audiography", Raven Press, 1986; PA1 2. R. D. Evans: "The Atomic Nucleus", Kreiger, 1955; PA1 3. J. C. Moyers: "A High Performance Detector Electronics System for Positron Emission Tomography", Masters Thesis, University of Tennessee, Knoxville, Tenn., 1990; PA1 4. U.S. Pat. No. 4,743,764 issued to M. E. Casey, et al, on May 10, 1988; PA1 5. R. A. DeKemp, et al.: "Attenuation Correction in PET Using Single Photon Transmission Measurement", Med. Phys., vol. 21, 771-8, 1994; PA1 6. S. R. Cherry, et al.: "3-D PET Using a Conventional Multislice Tomograph Without Septa", Jl. C. A. T., 15(4) 655-668. PA1 7. J. S. Karp, et al.: "Singles Transmission in Volume-Imaging PET With a .sup.137 Cs Source", Phys. Med. Biol. Vol. 40, 929-944 (1995). PA1 8. S. K. Yu, et al.: "Single-Photon Transmission Measurements in Positron Tomography Using .sup.137 Cs", Phys. Med. Biol. Vol. 40, 1255-1266 (1995). PA1 9. G. F. Knoll: Radiation Detection and Measurement, John Wiley & Sons (1989). PA1 10. S. R. Cherry, et al.: "Optical Fiber Readout of Scintilator Arrays using a Multi-Channel PMT: A High Resolution PET Detector for Animal Imaging", IEEE Transactions on Nuclear Science, Vol. 43, No. 3, 1932-1937 (June, 1996). PA1 11. J. A. McIntyre, et al.: "Construction of a Positron Emission Tomograph with 2.4 mm Detector", IEEE Transactions on Nuclear Science, Vol. 33, No. 1, 425-427 (February, 1986).
Both SPECT and CAT (or CT) systems are also well known to persons skilled in the art.
In order to achieve maximal quantitative measurement accuracy in tomography applications, an attenuation correction must be applied to the collected emission data. In a PET system, for example, this attenuation is dependent on both the total distance the two gamma rays must travel before striking the detector, and the density of the attenuating media in the path of travel. Depending on the location of the line of response within the patient's body, large variations in attenuating media cross section and density have to be traversed. If not corrected for, this attenuation causes unwanted spatial variations in the images that degrade the desired accuracy. As an example, for a cardiac study the attenuation is highest in the line of responses (LORs) passing through the width of the torso and arms, and attenuation is lowest in the LORs passing through from the front to the back of the chest.
Typically, the attenuation correction data in PET systems is produced by either: shape fitting and linear calculations using known attenuation constants, these being applicable to symmetric well-defined shapes such as the head and torso below the thorax (calculated attenuation); or through the measurement of the annihilation photon path's attenuation using a separate transmission scan (measured attenuation). The use of calculated attenuation correction, which introduces no statistical noise into the emission data, can be automated for simple geometries such as the head, and is the most prominent method used for brain studies. However, complexities in the attenuation media geometry within the chest have prevented the application of calculated attenuation from being practical for studies within this region of the body. Accordingly, transmission scanning has been utilized.
The total attenuation of a beam along a LOR through an object is equal to the attenuation that occurs for the two photons from an annihilation. Thus, the emission attenuation along the path can be measured by placing a source of gamma rays on the LOR outside of the body and measuring attenuation through the body along this line. It has been the practice to accomplish this attenuation measurement by placing a cylindrical positron emitter "sheet" within the PET tomograph's field of view (FOV) but outside of the region (the object) to be measured. The ratio of an already acquired blank scan (no object in the FOV) to the acquired transmission scan is calculated. These data represent the desired measured attenuation factors, which may vary spatially. These data are then applied to the emission data after a transmission scan of the object to correct for the spatial variations in attenuation.
There are two types of transmitter source units conventionally utilized in PET transmission scan data collection, both of which form a "sheet" of activity to surround the patient. One involves the placement of rings of activity aligned with detector rings around the inner face of the septa. The second type utilizes the rotation of one or more axially-oriented rods of activity in a circular path just inside the inner face of the septa.
The first of these two emitter systems (the ring source method) significantly reduces the sensitivity of the tomograph due to the close source-proximity dead time effects of the source activity on all of the detectors. Further, removal of this assembly is either performed manually by facility personnel or by a complex automated mechanical assembly. Large, cumbersome, out of the FOV shielding is required for storage of the automated source when not in use, adding to the depth of the tomograph tunnel and, thus increasing incidence of patient claustrophobia. The second type of emitter, using rotating source(s) suffers from the above-mentioned problems and also, due to the shielding requirements, reduces the patient tunnel diameter, further increasing patient claustrophobia symptoms.
Both of the above automated source transportation methods suffer from high mechanical component cost and from low sensitivity. Due to the dead-time-induced reduction in tomograph sensitivity, lengthy acquisitions are required in order to achieve usable low noise transmission scan data.
In order to reduce costs in scintillator detector applications, multiplexing techniques based on the use of fiber optics are advantageous. Those disclosures made by Cherry, et al. (Cherry), and McIntyre, et al. (Mcintyre), teach the use of fiber optics connected between the imaging detectors and multichannel photomultipliers (PMT's). Cherry discloses the use of a multi-channel PMT in association with an 8.times.8 array of bismuth germanate (BGO) crystals. As discussed by Cherry, a charge division readout board is used to convert the 64 signals into four position sensitive signals which determine the crystal interaction. In the earlier McIntyre article, the authors disclose the use of fiber optics coupled between the detectors and a number of multi-channel PMT's. Specifically, McIntyre teaches the use of 288 PMT's in association with 8,192 detectors, for reducing the number of required PMT's by a factor of about 28.4.
In the McIntyre embodiment, eight detector rings are each divided into four quadrants. Each ring is comprised of sixteen concentric rings. The respective quadrants for the eight detector rings are grouped together for a total of 256 detectors per quadrant group. Sixteen "coarse" fiber sets connect sixteen PMT's to the 256 detectors, with sixteen detectors in one ring quadrant connected to one PMT. Similarly, sixteen "fine" fiber sets connect sixteen PMT's to the 256 detectors, with corresponding detectors in each ring quadrant of a quadrant group being connected to one PMT. One PMT is connected to each ring quadrant. Thus, a total of 32 PMT's are required for determining the particular detector ".THETA." address within a quadrant. Similarly, 32 PMT's are required to determine the "r" address, corresponding to which of the concentric rings in a particular ring the detector is disposed. Finally, eight PMT's are required to determine which ring quadrant the detector is disposed. Thus, a total of 72 PMT's are required for each quadrant for a total of 288 PMT's in association with 8,192 detectors.
Therefore, it is an object of the present invention to provide a system for detecting coincident activity from a point source.
Another object of the present invention is to provide such a system which includes a detector dedicated to collecting attenuation data.
Yet another object of the present invention is to provide a system for detecting coincident activity while illuminating only a strip of the imaging detector in order to eliminate events not of interest in the attenuation measurement.
A further object of the present invention is to provide a collimated point source and dedicated detector whereby only a selected strip of the imaging detector is illuminated such that events unrelated to the attenuation are eliminated.
Still another object of the present invention is to provide an arrangement whereby gamma radiation detected by dedicated detectors is transmitted to a plurality of PMT's such that an address of each gamma radiation detector is readily determined and such that the total required number of PMT's is reduced relative to conventional devices.